Currently, the great majority of stents delivered transluminally and percutaneously to a stenting site in a human body are made of a biologically compatible material which is a metal. Many stents are made of stainless steel, and many others are made of nickel titanium shape memory alloy. The nickel titanium stents are invariably self-expanding stents that utilise a shape memory effect for moving between a radially compact transluminal delivery disposition and a radially larger stenting disposition after placement in the body. Stainless steel stents are often delivered on a balloon catheter, with inflation of the balloon causing plastic deformation of the material of the struts, but other stainless steel stents rely on the resilience of the steel to spring open when a surrounding sheath is retracted relative to the stent being deployed.
However, in all cases, it is difficult to endow the stent strut matrix with a degree of flexibility that comes anywhere near the degree of flexibility of the natural bodily tissue at the stenting site. The strength and resilience of the stent matrix, that serves to push radially outwardly the bodily tissue at the stenting site, is difficult to reconcile with the flexibility in bending that the natural tissue around the stent is capable of exhibiting, in normal life of the patient carrying the stent. It is one object of the present invention to improve the performance of a stent prosthesis in bending, after it has been deployed in the body of a patient.
To explain the problem, reference will now be made to applicant's WO 01/32102, specifically drawing FIGS. 3 and 4, and the text, of WO 01/32102. Indeed, accompanying drawing FIGS. 1 and 2 are the same as FIGS. 3 and 4 of WO 01/32102.
Looking at accompanying FIG. 1, we see part of the circumference of a tubular workpiece of nickel titanium shape memory alloy, in side view. The tube has a diameter D and a multiplicity of slits 20, 22 and 24, through the wall thickness of the tubular workpiece, all parallel to each other and to the longitudinal axis of the workpiece and creating out of the original solid tubular workpiece a lattice which can be expanded radially outwardly, (for example on a mandrel) to the expanded configuration of drawing FIG. 2 (again in side view). Out of the multitude of parallel slits can now be recognised as a sequence of 10 stenting rings, all displaying a zig-zag advance around the circumference of the prosthesis. Terminal zig-zag rings 30 are composed of 24 struts 32 interspersed by points of inflection 34, giving the end view of the prosthesis the appearance of a crown with twelve points.
The eight zig-zag rings at intermediate points along the length of the stent, between the two end rings 30, are referenced 36. They are made up of struts 38 which are all much the same length, somewhat shorter than end struts 32. Between any two struts of any of the zig-zag stenting rings there is a point of inflection 40. In the two end rings 30, all twelve of these points of inflection remote from the crown end of the terminal ring 30 are connected to a corresponding point of inflection 40, head to head, in the next adjacent internal stenting ring 36. However, between any two internal stenting rings 36, not all the twelve points of inflection, found spaced around the circumference of the prosthesis, are joined to corresponding points of inflection on the next adjacent stenting ring 36. Indeed, reverting to FIG. 1, it is easy to see that there will be only four connector portions 42, linking any two adjacent internal stenting rings 36.
Thinking about advance of the prosthesis of FIG. 1, in its compact disposition, along a tortuous, transluminal, delivery path to the stenting site, as the stent bends around a sharp bend in the delivery path, on the inside of any such bend, for example at point 44 on FIG. 1, the points of inflection facing each other across the gap 60 will approach one another. Depending on the length of the diametrically opposed connector portions 42 connecting stenting rings 36B and 36C, the two unconnected points of inflection will come into contact with each other in the middle of the gap 60, in dependence upon how sharp is the bend that the stent is negotiating in the tortuous path at that time. The longer the axial gap between adjacent stenting rings, the greater the capability of the stent for negotiating ever tighter bends in the delivery path lumen.
But what of the performance of the stent in bending, after it has been deployed at the stenting site.
We can see from FIG. 2 that the pattern of connector portions 42 is symmetrical. That is to say, standing on one of these connector portions, and looking along the length of the prosthesis, the pattern of connectors to the left of the line of view is a mirror image of the pattern of connectors to the right of that line of view. If we switch to consideration of drawing FIG. 3, which shows a portion of the strut network of the stent of FIGS. 1 and 2, this is more readily evident. Just as points of inflection on the inside of a tight bend of the stent in its compact disposition of FIG. 1 can butt up against each other face to face, so can the same phenomenon occur when the expanded stent of FIG. 2 is subject to sharp bending. Any such intermittent abutment of otherwise free points of inflection is liable to have negative effects including, for example, irritation or injury to bodily tissue caught between the abutting points of inflection, or even incipient buckling of the stent with the potential to reduce flow of bodily fluid through the stent lumen to dangerously low levels.
It is one object of the present invention to mitigate these risks.